High energy launch fiber terminations are widely used for laser surgical applications. Some of the fibers incorporate a taper; and some are straight fiber at the launch termination. These terminations are used, by way of example, in surgical lithotripsy. Typically, in such surgical fiber applications, the fibers, whether they incorporate a taper or are straight, are stripped of the buffer (polymer coating) and are fused with a quartz ferrule at the terminus. The terminus itself may be either mechanically or laser polished, and may incorporate an integral lens.
In conjunction with surgical endoscopes, it is necessary for the OD (outside diameter) of the fiber to be small so that it can easily pass through small working channels capable of use deep in the kidney. Typically, 200 .mu.m silica core, 220 .mu.m silica clad, 240 .mu.m polyimide buffer fiber has been employed. Since the cladding is a mere 10 .mu.m thick, its use at wavelengths longer than 2 .mu.m is questionable. A general "rule of thumb" for fiber is a cladding thickness which is five times the maximum wavelength of the light through the fiber. This rule is most important in applications where the fiber is highly stressed, that is bent at tight radii. In addition, the process of applying polyimide (solvent casting) causes the polyimide to be tight on the fiber. It basically shrinks into place in the final stages of its "cure". This shrinkage of the polyimide buffer also imparts stress to the relatively thin cladding. As a consequence, the rule mentioned above for cladding thickness is even more critical. It has been found that fibers of this type often burn up in surgery, sometimes damaging the endoscope.
The output or distal tip for endoscope fibers is prepared by cleaving or polishing (for larger core fibers). The first time, it is done in the factory or by the supplier; and then in each re-use the surgeon or nurse in the operating room cleaves the fiber. When polyimide buffer fiber is cleaved, the polyimide is stretched and a little flap of the polymer frequently remains extended beyond the end of the glass. This flap then is ignited on the first laser pulse, causing it to shrink back to the fiber face. Each subsequent pulse binds the residual carbon more tightly to the glass. The tip glows red; and the output is distorted. Finally, polyimide buffered fiber is prone to damage because of the relatively thin (typically 10 .mu.m thick) polymer coating. As a result, the fibers commonly break long before their usefulness is exhausted.
The focal spots of lasers used in medicine tend to be somewhat sloppy, as a result of poor maintenance, design of the launch, and laser modal instability. This causes the beam often to be presented larger than that which is specified by the manufacturer.
Where fiber terminations are glued to secure them in the ferrule, deposits of the outgassing adhesive (when heated by overfill energy) tend to contaminate the laser output lens. In some cases, the fiber termination explosively fails, firing shards of glass at the lens (or in some cases, window), destroying it. While this problem of outgassing adhesive, when it is struck by stray laser energy, is not particularly significant at low energy levels, the increased energy levels which continue to be applied in laser surgery greatly exacerbate the problem. Where a metallic beam block for dissipating stray laser energy is employed, the beam block typically is adhered to the end of the quartz ferrule with adhesive, which is subject to the same possibility of failure if that adhesive is contacted with sufficient laser energy.
In an effort to avoid the outgassing problems noted above for glued fiber terminations, another approach involves removing the polymer at the fiber tip and using some mechanism to prevent the adhesive (commonly, epoxy) from wicking all of the way to the front of the ferrule. This has been accomplished by using high viscosity adhesives, pre-heating the ferrules so the adhesive cures at a narrow bore before it can wick, or using a temporary adhesive to take up the space where the desired adhesive is not wanted. When a temporary adhesive is used, the desired adhesive is cured; and the temporary adhesive then is removed. All of these techniques are time consuming and produce low yield. They also do not work for PCS (plastic clad silica) fiber in that the polymer coating for PCS fiber also is cladding and must remain.
Another approach to remove adhesive from the immediate area of the fiber core at the laser focus involves countersinking of the ferrule at the fiber tip; so that the adhesive for securing the fiber to the ferrule is located a pre-established distance from the fiber tip. Countersinking also requires considerable labor and produces a low yield. In addition, it is not possible to countersink very far in most standard connectors, thereby requiring custom connectors. At the same time, it is still possible to have polymer at the fiber face, depending upon the manufacturer of the fiber.
Yet another attempt for eliminating the outgassing problem is to use a specifically modified connector tip to allow a circumferential crimp directly onto the fiber coating at the tip. To accomplish this it is necessary for the fiber to have polymer up to the fiber face, since crimping directly on glass is difficult and produces very low yields. The lowered mass of the connector (metal) at the fiber tip also can result in vaporization of the metal itself, which is even worse than polymer vapor. In addition, most surgical laser interlocks do not recognize the altered outside diameter of the connector which is required for the crimp.
It is desirable to provide a robust, small diameter fiber termination which overcomes the disadvantages of the prior art noted above, which is capable of handling high power inputs, which may undergo flash autoclave sterilization, and which does not employ adhesives in any area where stray energy can contact such adhesives.